Digital silicon photomultiplier for TOF-PET

ABSTRACT

A radiation detector includes an array of detector pixels each including an array of detector cells. Each detector cell includes a photodiode biased in a breakdown region and digital circuitry coupled with the photodiode and configured to output a first digital value in a quiescent state and a second digital value responsive to photon detection by the photodiode. Digital triggering circuitry is configured to output a trigger signal indicative of a start of an integration time period responsive to a selected number of one or more of the detector cells transitioning from the first digital value to the second digital value. Readout digital circuitry accumulates a count of a number of transitions of detector cells of the array of detector cells from the first digital state to the second digital state over the integration time period.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a Continuation of U.S. Ser. No. 15/000,336 filedJan. 19, 2016 which is a Divisional of U.S. Ser. No. 12/756,460 filedApr. 8, 2010 and since issued as U.S. Pat. No. 9,268,033 which is aDivisional of U.S. Ser. No. 11/467,670 filed Aug. 28, 2006 and sinceissued as U.S. Pat. No. 7,723,694 which is a Continuation of PCTapplication number PCT/IB2006/051089 filed Apr. 10, 2006 which claimsthe benefit of U.S. provisional application Ser. No. 60/674,034 filedApr. 22, 2005 and U.S. provisional application Ser. No. 60/682,246 filedMay 18, 2005, which are all incorporated herein by reference.

BACKGROUND

The following relates to the radiation detection arts. It particularlyrelates to high-speed radiation detectors for positron emissiontomography (PET), especially time-of-flight (TOF) PET, and will bedescribed with particular reference thereto. However, the followingrelates more generally to radiation detectors for single photon emissioncomputed tomography (SPECT), computed tomography (CT), and so forth, aswell as to high-speed radiation detectors for other applications such asastronomy.

In conventional PET, a radiopharmaceutical is administered to a humanpatient or other imaging subject. The radiopharmaceutical producesradiation decay events that emit positrons, which travel a very shortdistance before rapidly interacting with an electron of the surroundingimaging subject in an electron-positron annihilation event to producetwo oppositely directed gamma rays. The gamma rays are detected byradiation detectors surrounding the imaging subject as two substantiallysimultaneous radiation detection events that define a line of response(LOR) therebetween. Typically, the radiation detectors includescintillators that produce a burst or scintillation of light responsiveto each gamma ray detection, and an array of photomultiplier tubes(PMT's) optically coupled with the scintillators that convert the lightbursts into corresponding electrical signals. In some PET scanners, thePMT's are replaced by photodiodes that produce analog electricalcurrents proportional to the intensity of the light bursts.

Although the gamma rays are detected “substantially simultaneously”, ifone of the two involved radiation detectors is closer to theelectron-positron annihilation event than the other radiation detector,then there will be a small time difference between the two radiationdetection events. Since gamma rays travel at the speed of light, thistime difference between detections is typically around a few nanosecondsor less. In TOF-PET, the radiation detectors operate at sufficientlyhigh speed to enable measurement of this small time-of-flightdifference, which is then used to localize the electron-positronannihilation event along the LOR.

Accordingly, for TOF-PET the radiation detectors should havesub-nanosecond temporal resolution. PMTs are generally fast enough toperform TOF-PET imaging; however, PMTs are bulky, require high voltagebiasing, and are not well-suited for small pixel sizes desirable forhigh resolution. Conventional photodiodes are fast enough for TOF-PET,but lack internal amplification, leading to poor signal-to-noise ratios.To get sufficient signal with a conventional photodiode, acharge-sensitive amplifier is typically employed to integrate thesignal, which limits the bandwidth. Avalanche photodiodes can also beused; however, avalanche photodiodes typically suffer from high noiselevels and high temperature and bias sensitivity in the gain.

To address these difficulties, silicon photomultiplier (SiPM) detectorshave been proposed, for example in: E. A. Georgievskya et al., “Thesolid state silicon photomultiplier for a wide range of applications”,17^(th) Int'l Conf. on Photoelectronics and Night Vision Devices,Proceedings of SPIE vol. 5126 (2003); Golovin et al., “Novel type ofavalanche photodetector with Geiger mode operation”, Nuclear Instruments& Methods in Physical Research A, volume 518, pages 560-64 (2004). TheseSiPM detectors use a pixelated array of small avalanche photodiodesbiased in the breakdown region and interconnected in parallel. Theoutput is the analog sum of the currents of parallel-interconnectedavalanche photodiodes operating in limited Geiger-mode. Each detectedphoton in the SiPM detector adds on the order of 10⁶ electrons to theoutput current of the SiPM. The Geiger discharge responsive to photondetection is fast, providing sharp rising edges of the signal thatfacilitate precise time measurements. Energy- and temporal-resolutionscales with 1/sqrt(N) where N is the number of firing cells.

The SiPM device has certain disadvantages. The analog current producedby a photon detection is affected by bias voltage, operatingtemperature, and critical circuit parameters such as the quenchingresistance value. These factors can change the analog current producedby each photon detection, thus limiting the energy resolution of theSiPM. The analog configuration also has the disadvantages of producinghigh dark counts and allowing faulty avalanche photodiodes tosubstantially limit detector device manufacturing yield.

The following contemplates improved apparatuses and methods thatovercome the aforementioned limitations and others.

BRIEF SUMMARY

According to one aspect, a detector pixel is disclosed for use inconjunction with a scintillator that converts a radiation particle to aburst of light. An array of detector cells is provided. Each detectorcell includes a photodiode biased in a breakdown region and digitalcircuitry coupled with the photodiode. The digital circuitry isconfigured to output a first digital value in a quiescent state and asecond digital value responsive to detection of a photon by thephotodiode. Digital triggering circuitry is configured to output atrigger signal indicative of a start of an integration time periodresponsive to a selected number of one or more of the detector cellstransitioning from the first digital value to the second digital value.Readout digital circuitry accumulates a count of a number of transitionsof detector cells of the array of detector cells from the first digitalstate to the second digital state over the integration time period.

In some embodiments, digital timestamp circuitry is configured to outputa digital timestamp associated with the count. The digital timestamp isbased on a time of the trigger signal relative to a time referencesignal.

According to another aspect, a radiation detector includes ascintillator and an array of detector pixels as set forth in theprevious paragraph arranged to receive bursts of light produced by thescintillator in response to received radiation.

According to another aspect, a time-of-flight positron emissiontomography (TOF-PET) imaging system is disclosed. A plurality ofradiation detectors as set forth in the previous two paragraphs aredisposed to detect gamma rays emitted from an imaging region. Gamma raypair detection circuitry identifies two substantially simultaneous gammaray detections by two of the radiation detectors. A line of responseprocessor determines a spatial line of response connecting the two gammaray detections. A time of flight processor localizes a positron-electronannihilation event along the line of response based on a time differencebetween the two substantially simultaneous gamma ray detections.

According to another aspect, a method is performed in conjunction with ascintillator that converts a radiation particle to a burst of light.Digital circuitry is switched from a first digital value to a seconddigital value responsive to detection of a photon by a photodiode biasedin a breakdown region by the digital circuitry to define a switchingevent. A trigger signal indicative of a start of an integration timeperiod is generated responsive to a selected number of one or more saidswitching events associated with a plurality of said photodiodes. Acount of switching events associated with the plurality of saidphotodiodes is accumulated over the integration time period.

In some embodiments, the method further includes generating a digitaltimestamp associated with the accumulating over the integration timeperiod. The digital timestamp is based on a time of generation of thetrigger signal and a reference time signal.

According to another aspect, a radiation detector is disclosed,including a scintillator and circuitry for performing the method setforth in the previous paragraph.

One advantage resides in providing high data-rate radiation detectionfor TOF-PET, single photon emission computed tomography (SPECT),transmission computed tomography (CT), astronomy, and otherapplications.

Another advantage resides in providing a digital radiation detectoroutput.

Another advantage resides in providing a digitally timestamped detectoroutput.

Another advantage resides in providing improved spatial detectorresolution.

Another advantage resides in improved detector device manufacturingyield with low sensitivity to temperature, bias voltage, and processparameters.

Numerous additional advantages and benefits will become apparent tothose of ordinary skill in the art upon reading the following detaileddescription.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention may take form in various components and arrangements ofcomponents, and in various process operations and arrangements ofprocess operations. The drawings are only for the purpose ofillustrating preferred embodiments and are not to be construed aslimiting the invention.

FIG. 1 diagrammatically shows a TOF-PET system employing high-speedpixelated digital radiation detectors.

FIG. 2 diagrammatically shows a cross-sectional view of one of thepixelated digital radiation detectors of the TOF-PET system of FIG. 1.

FIG. 3 shows a general circuit diagram of one of the detector cells ofthe pixelated digital radiation detector.

FIG. 4A shows a more detailed circuit diagram of one embodiment of oneof the detector cells.

FIG. 4B shows a more detailed circuit diagram of another embodiment ofone of the detector cells.

FIG. 5 shows a circuit diagram of one pixel of the pixelated digitalradiation detector.

FIG. 6 shows a circuit diagram of one the pixelated digital radiationdetectors.

FIG. 7 shows a cross-sectional view of one physical layout embodiment ofthe pixelated digital radiation detector, in which the photodiodesdefine a photodiode layer and the digital circuitry is disposed in adigital circuitry layer separate from and electrically coupled with thephotodiode layer.

FIG. 8 shows a perspective view of another physical layout embodiment ofthe pixelated digital radiation detector, in which the photodiodesdefine a photodiode layer and the digital circuitry is disposed in thephotodiode layer interspersed amongst the photodiodes.

FIG. 9 shows a plan view of the light-sensitive area of a variant devicewhich includes the pixelated digital radiation detector area and anadditional proportional photodiode that produces an analog photocurrentwhen the flux of photons is high enough to saturate the pixelateddigital radiation detector area.

FIG. 10 shows an illustrative example of the measurement setup used inthe first stage of a defective cell disablement process for detectorsincluding analog circuitry.

FIG. 11 shows a block schematic of one imaging counter cell.

FIG. 12 shows a sensor block diagram.

FIG. 13 shows a photodetector incorporating a fuse for disablement.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

With reference to FIG. 1, a time-of-flight positron emission tomography(TOF-PET) scanner 8 includes a plurality of radiation detectors 10arranged to view an imaging region 12. In FIG. 1, the plurality ofradiation detectors 10 are arranged in several rings of detectors alongan axial direction; however, other arrangements of radiation detectorscan be used. Moreover, it is to be appreciated that the plurality ofradiation detectors 10 is diagrammatically illustrated; typically theradiation detectors are housed within a housing 14 of the scanner 8 andthus are not visible from the outside, and typically each ring ofradiation detectors includes hundreds or thousands of radiationdetectors. In some PET scanners, only a single ring of radiationdetectors is provided, in others, two, three, four, five, or more ringsof radiation detectors are provided. It should be appreciated thatdetector heads can be used in place of the detector ring structure shownin the Figures. The TOF-PET scanner 8 includes a couch 16 or othersupport for positioning a human patient or other imaging subject in theimaging region 12. Optionally, the couch 16 is linearly movable in theaxial direction generally transverse to the rings of radiation detectors10 to facilitate acquisition of three-dimensional imaging data.Additionally or alternatively, the imaging subject can be heldstationary, and the plurality of rings of radiation detectors used toacquire three-dimensional TOF-PET imaging data. In yet otherembodiments, only a single ring of detectors is provided, the imagingsubject remains stationary, and the resulting image is two-dimensional.

A suitable radiopharmaceutical is administered to the patient or otherimaging subject prior to initiation of TOF-PET imaging. Theradiopharmaceutical includes a radioactive substance that undergoesradioactive decay events that emit positrons. The positrons rapidlyannihilate with nearby electrons of the imaging subject. The resultingpositron-electron annihilation event produces two oppositely directedgamma rays having energies of 511 keV. The gamma rays travel at thespeed of light, i.e. ˜3×10⁸ meters/sec. Since the imaging region 12typically has a diameter or other characteristic dimension of about twometers or less, the time-of-flight for a gamma particle from theposition of the positron-electron annihilation event to one of thedetectors of the plurality of radiation detectors 10 is about a fewnanoseconds or less. Thus, the two oppositely directed gamma rays striketwo of the radiation detectors substantially simultaneously.

With continuing reference to FIG. 1 and with further reference to FIG.2, each radiation detector 10 includes a scintillator 20 that produces ascintillation or burst of light when the gamma ray strikes thescintillator 20. The burst of light is received by an array of detectorpixels 22 monolithically disposed on a silicon substrate 24. As will bedescribed, the detector pixels 22 are digital detector pixels thatoutput values including a digital representation of a count of photons(denoted “count” in FIG. 1) and a digital representation of a timestamp(denoted “timestamp” in FIG. 1) indicative of when the burst of lightcorresponding to the scintillation event was detected by the detectorpixel 22. Moreover, the plurality of radiation detectors 10 outputsindexing information including, for example, a detector index (denoted“n_(detector)” in FIG. 1) indicating which of the radiation detectors 10output the radiation detection event, and a detector pixel index(denoted “k_(pixel)” in FIG. 1) indicating which detector pixel orpixels of that radiation detector detected the burst of lightcorresponding to the radiation detection event. The scintillator 20 isselected to provide high stopping power for 511 keV gamma rays withrapid temporal decay of the scintillation burst. Some suitablescintillator materials are LSO, LYSO, MLS, LGSO, LaBr and mixturesthereof. It should be appreciated that other scintillator materials canalso be used. Although FIG. 2 shows the scintillator 20 as a singlecrystal, an array crystals can instead be used. Additionally, anoptional planar light pipe 26 can be interposed between the scintillator20 and the detector pixels 22 to improve transmission of photons of thescintillation light burst to the detector pixels 22. The scintillator 20and optional light pipe 26 are optionally encased in a reflectivecoating 28 which directs scintillation light toward the pixels 22.

With continuing reference to FIG. 1, the digital data concerningradiation detection events are processed by a pre-processor 30 thatperforms selected data processing. For example, if a scintillation eventis detected by a plurality of detector pixels, then the pre-processor 30can employ Anger logic or other processing to identify spatialcoordinates r for each radiation detection event and to estimate anenergy of the detected radiation particle. The resulting spatial andenergy information for each radiation detection event is stored in anevents buffer 32. In other embodiments, the scintillator layer isdivided into scintillator tiles sized to correspond with the detectorpixels, and each detector pixel is optically coupled with a singlescintillator tile. For example, each scintillator tile may include areflective coating similar to the reflective coating 28 to channel thescintillation photons to the coupled pixel.

A gamma ray pair detection circuitry 34 processes the radiationdetection events to identify pairs of substantially simultaneous gammaray detections belonging to corresponding electron-positron annihilationevents. This processing can include, for example, energy windowing (thatis, discarding radiation detection events outside of a selected energyfiltering window disposed about 511 keV) and coincidence-detectingcircuitry (that is, discarding radiation detection event pairstemporally separated from each other by greater than a selected timefiltering interval).

When a gamma ray pair is identified, a line-of-response (LOR) processor38 processes the spatial information pertaining to the two gamma raydetection events (for example, with the two events represented byspatial coordinates r₁ and r₂, respectively, as computed by thepre-processing 30) to identify a spatial line of response (LOR)connecting the two gamma ray detections. Since the two gamma raysemitted by a positron-electron annihilation event are oppositelyspatially directed, the electron-positron annihilation event is known tohave occurred somewhere on the LOR.

In TOF-PET, the radiation detectors 10 have sufficiently high temporalresolution to detect a time-of-flight difference between the two“substantially simultaneous” gamma ray detections. A time-of-flightprocessor 40 analyzes the time difference between the times (denoted“t₁” and “t₂” in FIG. 1) of the two gamma ray detection events tolocalize the positron-electron annihilation event along the LOR. Theresult, accumulated for a large number of positron-electron annihilationevents, is a set of histoprojections 42. A reconstruction processor 44reconstructs the set of histoprojections 42 into a reconstructed imageusing any suitable reconstruction algorithm, such as filteredbackprojection or iterative backprojection with correction. Theresulting reconstructed image is stored in an images memory 46, and canbe displayed on a user interface 48, printed, stored, communicated overan intranet or the Internet, or otherwise used. In the illustratedembodiment, the user interface 48 also enables a radiologist or otheruser to control the TOF-PET scanner 8; in other embodiments, a separatecontroller or control computer may be provided.

With reference to FIG. 3, each pixel 22 of the radiation detector 10includes an array of detector cells 50; FIG. 3 shows a general circuitdiagram of one such detector cell 50. A photodiode 52 is biased in abreakdown region and serves as an input to digitizing circuitry 54. Anoutput 56 of the digitizing circuitry 54 has a first digital valuecorresponding to a quiescent state, and transitions to a second digitalvalue responsive to detection of a photon by the photodiode 52. When thefirst photon of a scintillation burst is detected, the switching of theoutput 56 from the first digital value to the second digital valueactivates an open collector trigger line driver 60 which causes atrigger signal to be applied to a common trigger line or bus 62. Thetrigger signal in turn initiates a photon counter/FIFO buffer 66 (whereFIFO=“first in, first out”) that counts the switchings of the digitizingcircuitry 54 from the first digital value to the second digital valueover an integration time period started by the trigger signal. In someother embodiments, an acquisition enable line 67 initiates the photoncounter 66. A quenching circuit 70, which may be either active orpassive, limits current through the photodiode 52 and is configured tofacilitate transition the biasing circuitry from the second digitalvalue back to the first digital value. Thus, the detector cell 50 maycount more than one photon if the detector cell 50 is quenched back tothe quiescent first digital value before the integration time periodexpires. The final count stored in the photon counter/FIFO buffer 66 isaccessible via a digital bus 68.

The photodiode 52 is suitably biased in a Geiger-mode type of operation.When the photodiode 52 breaks down, large amount of charge (for example,about 10⁶ electrons per received detection in some photodiodes) isgenerated through the avalanche breakdown process. This charge istransported primarily through the quenching circuit 70, which has aneffective resistance of typically several hundred kilo-ohms to limit thecurrent flowing through the photodiode 52. With the current thuslimited, charge remaining in the photodiode 52 distributes spatially toreduce the electric field in the avalanche region of the photodiode 52.This screening quenches the avalanche process and causes remainingcarriers to be transported by drift out of the avalanche/depletion zone,causing recovery of the photodiode 52. Typically, the photodiode 52includes a guard ring (not shown) around the periphery that preventsavalanche breakdown at the edges of the photodiode 52. The guard ringstructure suitably acts like an ordinary reverse-biased PN diode withinternal fields too low for the avalanche breakdown to occur.

With reference to FIG. 4A, a more detailed circuit diagram of oneexample embodiment detector cell 50′ is described. This embodimentincludes a passive quenching circuit 70′ embodied as a quenchingresistor. On photon detection, the junction of the photodiode 52 breaksdown and an electric current starts to flow through the photodiode 52and the quenching resistor 70′. This current causes a voltage dropacross the resistor 70′, thus lowering the potential on the inverterinput. The voltage difference relative to VDD should be large enough todrive the inverter output into a “high” state. The switchingcharacteristics of the inverter can be optimized by adjusting thetransistor widths. The inverter output returns to a “low” stateautomatically when the photodiode 52 recovers from the breakdown.

With continuing reference to FIG. 4A, the detector cell 50′ furtherimplements inhibit logic 74 which does not switch off a faulty detectorcell completely, but rather prevents faulty detector cells fromgenerating false triggers. Faulty detector cells will generate excesscurrents which are taken into account in the trigger validation circuit(described later). When using the detector cells 50′, the trigger line62 is tied to a “high” level via a pull-up resistor (not shown in FIG.4A). This way, the triggers from all the detector cells 50′ arelogically “or”-ed together and the trigger line 62 is pulled down bythat detector cell which detects the first photon.

With reference to FIG. 4B, a more detailed circuit diagram of anotherexample embodiment detector cell 50″ is described, which includes anactive quenching circuit 70″ to speed up the discharge of the junctioncapacitance of the photodiode 52 to return to the quiescent level, thusreducing the recovery time of the photodiode 52. Shorter recovery timesare expected to lead to higher sensitivity, since a given detector cell50″ is more likely to count more than one photon during the integrationtime period when it recovers quickly, and are expected to lead to higherdynamic range and better energy resolution of the detector cell 50″. Thephoton counter 66 is enabled either by the trigger line 62 or adedicated line if a hierarchical trigger network is used, which ispulled down by that detector cell which detects the first photon, and isheld down by main pixel logic (not shown in FIG. 4B) for the integrationtime period. The number of detected photons accumulated by the photoncounter 66 is transferred from the photon counter 66 to a buffer orother digital storage (not shown in FIG. 4B) on the rising edge of thetrigger line 62 or a dedicated readout line. Subsequently, the counter66 is reset automatically, for example by the low level of the invertedand delayed signal on the trigger line 62, in preparation for the nextscintillation burst detection event. In this arrangement, the dead timebetween integration time periods can be as low as the buffer transfertime plus the reset time of the counter 66, which in some embodiments isexpected to be less than one nanosecond for CMOS implementations. Thedetector cell 50″ of FIG. 4B also includes inhibit logic 74 to preventfalse triggers from faulty detector cells.

With reference to FIG. 5, each pixel 22 of the radiation detector 10includes a two-dimensional array of detector cells 50 and associatedpixel-level digital pixel logic 80. Digital readout circuitry for thepixel 22 includes pixel-level digital acquisition and readout circuitry82 and associated circuitry at the detector cell level.

With reference to FIGS. 3 and 5, the digitizing circuitry 54 of eachdetector cell 50 provides a threshold-based binary digital outputindicating whether or not the photodiode 52 of that detector cells hasentered breakdown. The digital circuitry 54 outputs a first binary valuewhen the photodiode 52 is in its quiescent state, and transitions to asecond binary value when the photodiode current increases beyond athreshold value indicative of photon detection. The signal of eachphotodiode 52 is thus digitized at the level of the detector cell 50.Readout is performed by the pixel-level logic counting the digitaltransitions of the detector cells to produce a digital pixel outputindicative of the number of detected photons. Compared with the summingof analog photodiode currents to generate an analog pixel output as isdone in analog SiPMs, the digitize-and-count method of FIGS. 3 and 5 isfar less sensitive to bias variations, operating temperature variations,tolerances in components of the quenching circuit 70, or so forth. Aslong as these secondary effects do not cause erroneous switching ormissed switching of the thresholding digital circuitry 54, theygenerally do not affect the energy resolution of the detector cell 50.

In some readout approaches, the detector cells 50 are addressed like ina standard memory block, using address decoders for the rows and columnsof the array of detector cells 50. This solution provides a sequentialreadout of the cell data, in which case the pixel-level readoutcircuitry 82 can be a simple digital accumulator. In other readoutapproaches, cells lines are read out in parallel, with each line havingits own accumulator for the partial sum, and the partial sums are addedin a parallel adder tree. In yet other readout approaches, the addersare incorporated into the detection cells, so that the sum of the wholeline is obtained while clocking the data out and the line sum is readout from the last detector cell in the line. As the summation in thislatter readout approach can be pipelined, the readout architecture isfast, allowing short readout intervals.

If the detector cell-level photon counters 66 or counters of thepixel-level readout circuitry 82 are likely to saturate, then thecounters should not be allowed to wrap around. For example, a four-bitcounter counting from 0 . . . 15 should not be allowed to increment from15 back to 0. By avoiding wrap-around, saturation of the pixel 22 can bedetected when the counter reads its highest value (e.g., 15 for afour-bit counter). The number of bits for avoiding wrap-around dependssolely on the minimum anticipated cell recovery time and the maximumlength of the integration period. While the integration window is adesign parameter, the cell recovery time is of a statistical nature, asphoton detection probability is a function of the slowly risingover-voltage during cell recovery. In an actively quenched cell however,a minimum recovery time is defined by the monoflop delay. Thus, in thiscase, it is possible to design the counter wide enough to avoidoverflow. The digital bus 68 can be either a parallel or a serial bus,depending on space and time constraints.

With continuing reference to FIG. 5, the digital pixel logic 80 furtherincludes trigger digital circuitry 84, trigger validation circuitry 85,and an output buffer 86 that stores the photon count of the pixel 22.The trigger digital circuitry 84 accesses a reference clock 88 (shown asan electrical trace that is connected to a suitable oscillator or otherclocking device not shown in FIG. 5) to provide a time reference for thetrigger digital circuitry 84. The trigger digital circuitry 84determines the time stamp of a radiation detection event in a global(for example, scanner) time frame. The trigger digital circuitry modules84 of all the pixels 22 of the scanner run synchronously at a precisionof preferably less than 100 ps. The reference signal 88 is used tosynchronize the trigger digital circuitry modules 84 of the pixels,providing them with a common time base for the entire scanner. In someembodiments, the integration time period is a fixed time intervalstarting at the occurrence of the trigger signal. In other embodiments,the integration time period is dynamically terminated when the rate ofnew counts decreases below a threshold value.

The trigger digital circuitry 84 is also preferably configured to outputthe digital timestamp (see FIG. 1) associated with the count. Thedigital timestamp is based on a time of the trigger signal output by thetrigger line driver 60 of the first one of the detector cells 50 thatdetects a photon from a scintillation burst. The pixel logic 80optionally still further includes data correction registers and inhibitsequence drivers. Automated test and calibration circuitry 87 is alsooptionally implemented by the pixel logic 80. In one test/calibrationmethod, the dark count rate of the pixel 22 (possibly includingbackground counts produced by intrinsic radioactivity of thescintillator 20) is monitored. In another test/calibration method, anexternal excitation from a test charge injected into the detector cells50 is used to test and calibrate the pixel 22.

With continuing reference to FIG. 5, it will be appreciated that due todark currents, crosstalk, thermal excitations, or so forth, it ispossible that one of the detector cells 50 may produce an inadvertenttrigger signal starting an integration time period. The triggervalidation circuitry 85 validates the trigger signal and aborts theintegration if it is determined that the trigger signal was false. Inone approach, the trigger validation circuitry 85 analyzes the currentflowing through the bias network of the pixel 22. If the total currentstays below a certain current threshold for a selected time interval(e.g. for 10 nanoseconds into the acquisition time period) as measuredby a discriminator or other circuitry, then the acquisition is abortedand an automatic reset sequence is initiated in preparation for the nexttrigger. If the current exceeds the current threshold, the discriminatoroutput will rise to a ‘high’ level and the acquisition will continue. Insome embodiments, rather than using a fixed integration time period, thefalling edge of the bias current discriminator is used to detect the endof the scintillation burst so as to adapt the integration time period tosubstantially match the end of the acquisition interval. This cansuppress pile-up in high count rate applications. Another suitablemethod makes use of the fact that the probability of two thermallygenerated triggers inside a short time window decreases with thedistance of the triggering cells since thermal triggers are generallynot correlated. In contrast, the scintillation burst should act ondetector cells 50 across the light-sensitive area of the pixel 22. Thus,the trigger validation circuitry 85 can analyze the triggers fromindividual detector cells 50, for example, and validate the triggersignal if two distant lines generate a trigger signal within a selectedtime window. Other approaches for trigger validation can also be used,such as employing a current sensor with adjustable discriminator set ata trigger threshold higher than the single photon level.

In some other embodiments, the counter 66 is triggered by theacquisition enable line 67. Triggering on the first photon can beproblematic if there is a high background flux of photons unrelated topositron-electron annihilation events. This background can be the resultof, for example, a secondary slow decay mode of the scintillator. Insuch cases, detector cells fire frequently, increasing the dead time ofthe pixel. To provide more robust counter initiating, at the detectorcell level (FIG. 3, 4A, or 4B) the photon counter is enabled by theseparate, ‘acquisition enable’ line 67 which is pulled down by the pixellogic on either the detection of the first photon (trigger line goesdown) or by the discriminator of the trigger validation circuit 85 whenthe current through the bias network has exceeded the user-definedtrigger level. This line defines the length of the integration windowand is driven by the pixel logic. At the detector pixel level (FIG. 5),the trigger validation circuit 85 is extended to include a multiplexer89 selecting either the trigger line 62 (for a single photon trigger) orthe leading edge discriminator output (for multiple photon trigger) asthe input to the time to digital converter/trigger validation circuits.The trigger validation circuit 85 is extended to provide the‘acquisition enable’ signal 67 to the detector cells 50, 50′,50″.

Alternatively, if triggering at single-photon level is not required, asuitable logic can be implemented to generate the trigger signal if aselected number of cells (trigger lines) become active at the same time.This implementation has the practical advantage requiring only digitalcomponents. However, in this case, the threshold is defined onlystatistically. In some other embodiments, the open collector driver isoptionally omitted from the detector cells and a modified design is usedin the trigger validation circuit.

With continuing reference to FIG. 5 and with further reference to FIG.6, the pixels 22 are arranged in a two-dimensional array to define thelight-sensitive surface of the pixelated radiation detector 10. Theembodiment shown in FIG. 6 uses a pixel readout in which each line ofpixels 22 is read out by FIFO buffers 90. The output buffers 90 eachinclude tristate output buffers allowing the data to be transferred overa shared digital data bus 92. Optionally, the events are sortedaccording to their time stamps by the readout arbitration in the lineoutput buffers 90 and also by the shared bus arbitration by outputbuffer 94, thus leading to a stream of event data which is sorted overtime. This optional feature substantially simplifies the search forcoincident events. A data request daisy-chain is suitably used for writeaccess arbitration. The daisy-chained sums are transferred to theradiation detector output buffer 94 for transfer off-chip.

With reference to FIGS. 7 and 8, in some embodiments the digitalcircuitry (such as the digital biasing circuitry 54, 54′,54″, digitaltriggering circuitry 60, 60′, 60″, 84, and readout digital circuitry 66,82) of the radiation detector 10 are defined by CMOS circuitry disposedon the silicon substrate 24. Various physical layouts can be used. In avertically segregated layout shown in FIG. 7, the photodiodes 52 of thearray of detector cells 50, 50′, 50″ define a photodiode layer 100, andthe digital circuitry are disposed in a CMOS digital circuitry layer 102separate from and electrically coupled with the photodiode layer 100. Inan alternative layout shown in FIG. 8, the photodiodes 52 define aphotodiode layer 100′, and the CMOS digital circuitry (such as thedigital biasing circuitry 54, 54′, 54″, digital triggering circuitry 60,60′, 60″, 84, and readout digital circuitry 66, 82) are disposed in thephotodiode layer 100′ interspersed amongst the photodiodes 52.

Because CMOS logic draws power only when switching states, only thoseparts of the radiation detector 10 that are continuously activelyclocked by the clock 88 will contribute to the baseline powerconsumption. Since the pixel 22 is activated by a trigger signalgenerated by one of the photodiodes 52 which are biased in the breakdownregion in the quiescent state, power consumption is dependent on thephoton detection rate and, thus, on the flux of received photons plusthe dark count rate. Control of power consumption of the pixel 22 can beimplemented by deliberately increasing the dead time of an individualpixel between two acquisitions. This could be done automatically by thepixel logic 80 depending on the temperature of the pixel. Thetemperature of the pixel can be measured directly by a temperaturesensor (not shown) or estimated indirectly from the dark count rate ofthe pixel 22.

Since CMOS logic draws power only when switching states, the overallpower consumption can be dramatically reduced by using a CMOSimplementation over an analog implementation. For example, in someembodiments of an analog implementation, the power consumption perchannel is 30 mW and the global part of the chip is 162 mW. For a morepractical implementation, such as on a clinical apparatus with 28,336channels or 1890 chips, the power consumption would be a constant 1156W. On the other hand, the power consumption for a CMOS implementation,such as the various implementations described herein, has two differentvalues, a static value and a dynamic value. The static power consumptionis the power required when there are no counts and hence no switching ofstates. It does include power for the logic of for the dynamic switchingas the logic must be ready to receive counts. The dynamic powerconsumption is the power required when the detector is activelyreceiving counts, and hence switching states. The power consumption inactive state is dependent on the amount of activity; the more counts andswitching of states, the power that is required. The static powerconsumption for a similar 1890 chip detector is about 10 W or less. Thedynamic power consumption can vary, depending on the activity, but istypically about 300 W or less.

A problem can arise if the scintillation burst of light produces a fluxof photons that is high enough to cause substantially all of thedetector cells 50, 50′, 50″ of one or more of the pixels 22 totransition from the first digital state to the second digital stateduring the integration time period. In this case, the pixel 22saturates, and the actual intensity (that is, the flux of photons) isnot accurately measured. This saturation problem can be addressed invarious ways.

In one approach, the photosensitive area defined by the photodiodes 52is broken into a larger number of smaller photodiodes. The reduced areaof each photodiode reduces the likelihood that it will detect a photon.The larger total number of photodiodes provides higher pixel-levelsensitivity to the flux of photons, although it generally does not fullycompensate for the reduced area of each cell. The detector cells shouldhave some separation to reduce optical crosstalk between neighboringdetector cells. Typically, the separation of the cells is in the orderof several microns or less, when trenches filled with opaque materialare used for the separation. Thus, increasing the number of cellsgenerally reduces the ratio of sensitive area to the total area of thecell to some degree. Additionally, increasing the number of detectorcells, while keeping the cell size constant, typically leads to aproportional increase of the dark count rate.

With reference to FIG. 9, in another approach for addressing thesaturation problem, a proportional photodiode 110 is included in thephotosensitive area. The proportional photodiode 110 is larger than thephotodiodes 52 used in digital detection. The proportional photodiode110 is configured to produce an analog photocurrent proportional to theflux of photons impinging upon the pixel 22 when said flux of photons ishigh enough to cause substantially all of the detector cells 50, 50′,50″ of the pixel 22 to transition from the first digital state to thesecond digital state during the integration time period. Although shownalong one side of the array of pixels 22 for simplicity of fabrication,the proportional photodiode 110 can be located in other positionsrespective to the array, such as centered in the array or at a corner ofthe array. Moreover, in some embodiments the proportional photodiode 110may be distributed as a plurality of smaller electrically interconnectedproportional photodiodes, such as a proportional photodiode located ateach corner of the array of pixels 22. In the variation of FIG. 9, thetrigger signal output by the first one of the photodiodes 52 to detect aphoton is still suitably used to provide the timing information for thegamma ray detection event. Thus, the timestamp output by the radiationdetector 10 is used; however, if the digital photodiodes 52 saturate,then the photocurrent produced by the proportional photodiode 110 isused to indicate photon flux intensity rather than using the digitalcount. The proportional photodiode 110 can be a conventional PIN diode,an avalanche photodiode with integrated analog or digital readoutcircuitry, or the like.

The pixelated digital radiation detectors are described herein inconjunction with an example TOF-PET application. However, the skilledartisan can readily adapt the disclosed pixelated digital radiationdetectors for other application, such as single-photon emission computedtomography (SPECT) imaging, transmission computed tomography (CT)imaging, astronomy applications, and so forth. For radiation detectionapplications in which the photodiodes 52 are directly sensitive to theradiation, the scintillator 20 is suitably omitted from the radiationdetector 10.

One skilled in the art should understand that while most of theembodiments have been described in conjunction with digital circuitry,portions of the invention can be implemented in conjunction with analogcircuitry. For example, the following description provides a method ofdisabling defective cells in an analog circuitry system. Suchembodiments are incorporated within the scope of this disclosure.

A defective cell disabling method for an analog circuit system cancomprise of two separate stages, namely a sensing stage and acalibration stage. During the sensing stage, a SiPM array or deviceunder test (DUT) is biased at the nominal bias voltage above thresholdin a light-tight setup. The Geiger-discharge in semiconductors generatessecondary light photons, approximately 3 per 100,000 electrons in thejunction on average. Thus, a cell with gain 1,000,000 will generateabout 30 optical photons. The average wavelength of these photons isabout 1 μm, thereby enabling the photons to travel large distances insilicon before being absorbed. Some of these photons trigger breakdownsin neighboring cells, commonly referred to as optical crosstalk, ifproper shielding is not used. Other photons can escape the silicon andcan be detected by appropriate single photon detectors. The sensingdetectors must be 1:1 coupled to the DUT cells. Thus, the trigger rateof the sensing detectors can then be directly associated with the darkcount rate of individual cells. Additional measurement of the chargepulse of the DUT can be used to directly measure the gain and itsvariation for individual DUT cells. However to collect sufficientstatistics, such measurement would likely mean a significant increase ofthe measurement time.

Based on the data acquired in the sensing stage, a laser beam willdisable faulty cells. Additionally, the number of active cells per pixelcan be adjusted to equalize the dynamic range of the pixels, ifrequired. In some implementations, a fuse is used to disable the faultycells. While a fuse would undesirably consume additional area, this canbe minimized if the fuse is placed over the guard ring. Anotheralternative would be to cut the poly resistor itself.

An illustrative example of the measurement setup used in the first stageis shown in FIG. 10. In FIG. 10, a single photon counter array 200 is1:1 coupled to the DUT 210 using a collimator structure 220. One skilledin the art should understand that if the sensing detector has the samepixel size as the DUT, proximity coupling could be used to increase thesensitivity of the system. The single photon counter array 200 must havesignificantly lower dark count rate and thus has to be cooled down to atleast −50° C. Each detector 230 in the photon counter array 200 istriggered by photons emitted by the Geiger-mode discharge. The detectorindicates the event by pulling down the row and column lines andstarting a hold-off interval to avoid double counting of the same event.The length of the hold-off interval must be adjusted to the recoverytime of the DUT. An active quenching/recharge circuit 240 can be used toobtain well-defined hold-off intervals. Additional circuits can be usedto measure the charge of the pulse in correlation to the coordinates ofthe event. A block schematic of one imaging counter cell is shown inFIG. 11, while a sensor block diagram is shown in FIG. 12.

Increasing the DUT temperature can be used to accelerate themeasurement. In the calibration stage, the pixel dark count rate andgain data is used to select a subset of cells that will be disabled.This can be any number of defective cells as well as other cells thatcan be disabled to provide uniformity. To achieve this, a laser is usedto cut the fuses in these cells, as illustrated in the modified detectorcell shown in FIG. 13.

Regardless of whether a digital or analog disablement process is used, areport can be generated allowing a user to determine how many cells weredisabled because they were deemed faulty. The report can further providethe location of the disabled faulty cells. The location of the disabledfaulty cells can, in some embodiments, be used to disable other cells.Typically this would be done in some sort of geometrical pattern toallow for more uniform detection of radiation about the detector.Furthermore, the disablement of other cells can be automatic, inresponse to manual input or feedback, or a combination thereof.

The resulting silicon photomultiplier array will have lower dark countrate at the expense of decreased sensitivity because of the area lostdue to dead cells. The loss in dynamic range can be accounted forbeforehand by integrating higher number of smaller-sized cells in thepixels. It should also be appreciated that the fuse implementation canbe used in combination with digital circuitry. For example, the fuse canbe used for calibration, while the digital circuitry is used for thecount detection. Other embodiments incorporating these types of ideasare also contemplated by this disclosure.

In some embodiments in which a trigger at the single photon level is notneeded, a leading edge discriminator can be used to generate the triggersignal and to suppress dark counts. In other embodiments the triggersignal can be generated digitally by applying a logical operation on thetrigger lines. For example, a pixel can be subdivided into two halves,or blocks, and the trigger signal is only generated if both halvesdetect the photon. In such embodiments, the number and size of theblocks can be adjusted to set the average threshold and the selectivity.Of course, other similar designs can be implemented, including, but notlimited to, other geometries and other ways of correlating pixel blocks.

The invention has been described with reference to the preferredembodiments. Obviously, modifications and alterations will occur toothers upon reading and understanding the preceding detaileddescription. It is intended that the invention be construed as includingall such modifications and alterations insofar as they come within thescope of the appended claims or the equivalents thereof.

The invention claimed is:
 1. A radiation detector comprising: ascintillator; and an array of detector pixels disposed monolithically ona common silicon substrate and arranged to receive bursts of lightproduced by the scintillator in response to received radiation, whereineach detector pixel includes: an array of detector cells, each detectorcell including a photodiode biased in a breakdown region and digitalcircuitry coupled with the photodiode, the digital circuitry beingconfigured to output a first digital value in a quiescent state and asecond digital value responsive to detection of a photon by thephotodiode; digital triggering circuitry configured to output a triggersignal indicative of a start of an integration time period responsive toa selected number of one or more of the detector cells transitioningfrom the first digital value to the second digital value; and readoutdigital circuitry that accumulates a count of a number of transitions ofdetector cells of the array of detector cells from outputting the firstdigital value to outputting the second digital value over theintegration time period.
 2. The radiation detector as set forth in claim1 wherein (i) the photodiodes of the array of detector cells define aphotodiode layer, and (ii) the digital circuitry, digital triggeringcircuitry, and readout digital circuitry define a digital circuitrylayer separate from and electrically coupled with the photodiode layer.3. The radiation detector as set forth in claim 1 wherein (i) thephotodiodes of the array of detector cells define a photodiode layer,and (ii) the digital circuitry, digital triggering circuitry, andreadout digital circuitry are disposed in the photodiode layerinterspersed amongst the photodiodes.
 4. The radiation detector as setforth in claim 1 wherein the digital circuitry, digital triggeringcircuitry, and readout digital circuitry are CMOS circuitry.
 5. Theradiation detector as set forth in claim 1 further including:multiplexing circuitry also disposed monolithically on the commonsilicon substrate, the multiplexing circuitry digitally multiplexingcounts produced by the readout digital circuitry of the detector pixelsto generate a digital radiation detector output signal.
 6. The radiationdetector as set forth in claim 1 wherein the detector pixels do notinclude analog photodiode current summing circuitry.
 7. The radiationdetector as set forth in claim 1 further comprising: a planar light pipeinterposed between the scintillator and the array of detector pixels. 8.The radiation detector as set forth in claim 7 further comprising: areflective coating encasing the scintillator and the planar light pipeand arranged to direct scintillation light toward the array of detectorpixels.
 9. The radiation detector as set forth in claim 1 wherein thescintillator is divided into scintillator tiles with each detector pixeloptically coupled with a single scintillator tile.
 10. The radiationdetector as set forth in claim 9 wherein each scintillator tile includesa reflective coating arranged to channel scintillation photons to theoptically coupled single detector pixel.
 11. A positron emissiontomography (PET) scanner comprising: a plurality of radiation detectorsas set forth in claim 1 arranged as one or more rings of detectorssurrounding an imaging region; and pair detection circuitry configuredto process radiation detection events from the plurality of radiationdetectors using an energy filtering window and a time filtering intervalto identify pairs of substantially simultaneous 511 keV gamma raydetections.
 12. A positron emission tomography (PET) scanner comprising:a plurality of radiation detectors arranged as one or more rings ofdetectors surrounding an imaging region and comprising scintillators anddigital detector pixels arranged to detect scintillation events producedby the scintillators in response to gamma rays striking thescintillators, wherein each digital detector pixel includes: an array ofdetector cells, each detector cell including a photodiode biased in abreakdown region; and digital triggering and readout circuitryconfigured to output a digital representation of a count of photonsdetected by the array of detector cells in response to the scintillationevent and a digital representation of a timestamp indicative of when thescintillation event was detected by the digital detector pixel; and pairdetection circuitry configured to process detected scintillation eventsfrom the plurality of radiation detectors using energy filtering andtime filtering to identify pairs of substantially simultaneous 511 keVgamma ray detections.
 13. The PET scanner of claim 12 wherein thescintillator is divided into scintillator tiles with each digitaldetector pixel optically coupled with a single scintillator tile. 14.The PET scanner of claim 13 wherein each scintillator tile includes areflective coating arranged to channel scintillation photons to theoptically coupled single digital detector pixel.
 15. The PET scanner ofclaim 14 wherein: each digital detector pixel further includes a planarlight pipe interposed between the scintillator tile and the opticallycoupled single digital detector pixel, and the reflective coatingencases the scintillator tile and the planar light pipe to channelscintillation photons to the optically coupled single digital detectorpixel.
 16. The PET scanner of claim 12 wherein the digital detectorpixels comprise CMOS digital detector pixels disposed on a siliconsubstrate.
 17. The PET scanner of claim 16 wherein: the photodiodes ofthe array of detector cells define a photodiode layer, and the digitaltriggering and readout circuitry define a digital circuitry layerseparate from and electrically coupled with the photodiode layer. 18.The PET scanner of claim 16 wherein: the photodiodes of the array ofdetector cells define a photodiode layer, and the digital triggering andreadout circuitry is disposed in the photodiode layer interspersedamongst the photodiodes.
 19. The PET scanner of claim 12 wherein thedigital detector pixels do not include analog photodiode current summingcircuitry.
 20. A positron emission tomography (PET) scanner comprising:a plurality of radiation detectors arranged as one or more rings ofdetectors surrounding an imaging region and comprising scintillators anddigital detector pixels arranged to detect scintillation events producedby the scintillators in response to gamma rays striking thescintillators, wherein the digital detector pixels are configured tooutput digital representations of counts of photons detected by thedigital detector pixels in response to scintillation events and digitalrepresentations of timestamps indicative of when scintillation eventswere detected by the digital detector pixels; and pair detectioncircuitry configured to process detected scintillation events from theplurality of radiation detectors using energy filtering and timefiltering to identify pairs of substantially simultaneous 511 keV gammaray detections.
 21. The PET scanner of claim 20 wherein the digitaldetector pixels do not include analog photodiode current summingcircuitry.
 22. The PET scanner of claim 20 wherein the scintillators aredivided into scintillator tiles each optically coupled with a singledigital detector pixel and having a reflective coating encasing thescintillator tile to channel scintillation photons to the opticallycoupled single digital detector pixel.